The present invention relates to a magnetic resonance imaging method for obtaining slice images using NMR (nuclear magnetic resonance), ESR (electron spin resonance) and the like.
The present invention relates, more particularly, to an improvement for restraining artifacts due to moving parts in a sample, e.g., a patient.
The present invention also relates to reduction of artifacts due to body movement of the sample.
The present invention also relates to a magnetic resonance imaging method for imaging the movement such as flow velocity and flow rate.
FIG. 1 is a block diagram showing a typical magnetic resonance (MR) imaging system. As illustrated, it comprises a static magnetic field generator 2 for applying a static magnetic field in the Z-axis to a sample 1, such as a patient to be examined, and a diagnosis table 3 for moving the sample into the static magnetic field generator 2.
The imaging system further comprises an RF (radio frequency) coil 4 for applying an RF energy such as an RF magnetic pulse to the sample 1 and receiving an MR signal from the sample 1, a matching device 5 which converts the impedance of the RF coil 4, a transmit/receive switch 6 connected to the matching device 5, a transmitter 7 for applying an RF energy to the sample 1 through the transmit/receive switch 6 and the RF coil 4, and a receiver 8 for receiving the MR signal from the sample 1 through the RF coil 4 and the transmit-receive switch 6.
The imaging system further comprises an X-axis gradient magnetic-field coil 9 for applying a magnetic field pulse inclined in the direction of the X-axis, an X-axis gradient magnetic-field power supply 10 for energizing the X-axis gradient magnetic-field coil 9, a Y-axis gradient magnetic-field coil 11 for applying a magnetic field pulse inclined in the direction of the Y-axis, a Y-axis gradient magnetic-field power supply 12 for energizing the Y-axis gradient magnetic-field coil 11, a Z-axis gradient magnetic-field coil 13 for applying a magnetic field pulse inclined in the direction of the Z-axis, and a Z-axis gradient magnetic-field power supply 14 for energizing the Z-axis gradient magnetic-field coil 13.
A sequence controller 15 is provided for controlling the transmitter 7, the receiver 8 and the gradient magnetic field power supplies 10, 12 and 14 with a predetermined sequence, and controlling various operations of the entire system, such as irradiation of the RF pulses, reception of the MR signals, generation of the gradient magnetic fields Gs, Gr and Ge, and operation of the diagnosis table 3.
A computer 16 is connected to the the receiver 8 and the sequence controller 15 and generates control data for generation of the image construction, and performs processing of the MR signals from the receiver 8 and processing for the image construction by means of, for example, two-dimensional Fourier transformation.
A console panel 17 serves as an input/output terminal for manually inputting, into the computer 16, commands and data including parameters necessary for the image construction.
An image display unit 18 is connected to the console panel 17 for displaying the generated slice images and other data.
The RF coil 4 irradiates RF pulses which are generated at the transmitter 7, and the RF pulses are transmitted via the transmit/receive switch 6 and the matching device 5. The MR signals from the sample 1 are received by the RF coil 4 and transmitted via the matching device 5 and the transmit/receive switch 6 to the receive 8.
The operation of the prior art MR imaging method will now be described with reference to the pulse sequence chart of FIG. 2. In this method, the gradient field echo method is used in combination with the two-dimensional Fourier transformation (2DFT) method. The magnetic field in the direction of the Z-axis (generated by the Z-axis magnetic field coil 13) is the slicing magnetic field Gs for designating the slice plane. The magnetic field in the direction of the X-axis (generated by the X-axis gradient magnetic field coil 9) is the signal reading magnetic field Gr for collecting signals. The magnetic field in the direction of the Y-axis (generated by the Y-axis gradient magnetic field coil 11) is the phase encoding magnetic field Ge. Moreover, the RF magnetic field pulse RF is a .alpha..sup.o (.alpha..ltoreq.90) pulse.
First, a sample 1 is inserted in the static magnetic field generator 2, the RF coil 4, and the gradient magnetic field coils 9, 11 and 13. Then the RF coil 4 and the Z-axis gradient magnetic field 13 are driven by the sequence controller 15 to apply an RF magnetic field pulse RF having a selected frequency. At the same time, a slice magnetic field Gs1 for designating the slice plane is applied to the sample 1. As a result, energy is imparted to the nuclear spins in the desired slice plane within the sample. The RF magnetic field pulse RF assumes its peak value when the pulse area of the slice magnetic field Gs1 becomes 1/2 (of its full area).
Then, a slice magnetic field Gs2 having an opposite polarity and having a pulse area (hatched) equal to 1/2 of the slice magnetic field Gs1 (i.e., the area B2 shown by hatching is equal to the area B1 also shown by hatching) is applied to rephase the magnetic spins (in the direction of the slice i.e., the Z-axis). At the same time, Y-axis gradient magnetic field coil 11 and the X-axis gradient magnetic-field coil 9 are driven to apply the phase encoding magnetic field Ge and the signal reading magnetic field Gr1 for the time T. The reading magnetic field Gr1 is applied for dephasing the MR signals in the direction of the X-axis. By the application of the signal reading magnetic field Gr1, the MR signal R is phase-disordered.
Subsequently, while a signal reading magnetic field Gr2 having an opposite polarity is applied, the MR signal R due to the gradient magnetic field echo is received by the RF coil 4, and supplied to the computer 16. The MR signal R which has been dephased is then rephased gradually by the application of the signal reading magnetic field Gr2, and thereafter dephased again. The time-point at which the MR signal R assumes its peak is dependent on the time-points at which the RF magnetic field pulse and the signal reading magnetic field are applied, and is dependent upon expiration of the echo time TE after the peak of the RF magnetic field pulse, and when the hatched pulse area of the signal reading magnetic field Gr2 equal the pulse area of the signal reading magnetic field Gr1.
The MR signals R are collected at predetermined sampling points for the period (t+t') of application of the signal reading magnetic field Gr2.
The above sequence has a repetition period TR corresponding to the image contrast. The above sequence is repeated the number of times corresponding to the predetermined number of pixels N, with the pulse of the phase encoding magnetic field Ge having its area (or, in the embodiment illustrated, the intensity if the pulse width is fixed) varied as indicated by the broken lines, and a plurality of MR signals R are sampled and received as a series of pulses. As a result, resolution of the MR signals R in the Y-axis can be made. For instance, if the number of pixels is 256.times.256, the number of samples received in one cycle is not less than 256 and the number of repetitions (signal collections) will be 256.
The computer 16 performs the two-dimensional Fourier transformation on the MR signals to reconstruct the images of the slice planes having a desired matrix size of N.times.N. The reconstructed image is displayed on the display unit 18.
On the assumption that the MR signals R are received from a stationary part and a moving part in the sample 1, and the time of commencement of the application of the signal reading magnetic field Gr1 is assumed to be t=0, the phase deviations .phi..sub.0 and .phi..sub.1 of the MR signals R from the stationary part and the moving part at the time t=2T will be considered.
If -P and P represent the intensities of the signal reading magnetic fields Gr1 and Gr2, r the gyromagnetic ratio, and X.sub.0 the coordinate of the stationary part, the phase deviation .phi..sub.0 of the stationary part during t=0 to 2T is given by the sum of the integral of the signal reading magnetic field Gr1 over t=0 to T: ##EQU1## and the integral of the signal reading magnetic field Gr2 over time t=T to 2T: ##EQU2## That is ##EQU3## This means that the phase deviation .phi..sub.0 of the MR signals R from the stationary part becomes zero at t=2T, and the disorder of the phase is eliminated.
On the other hand, if the coordinate of the moving part such as blood flow which moves at a constant velocity v is represented by (X.sub.1 +vt), the phase deviation .phi..sub.1 is given by: ##EQU4## Thus, at t=2T, the phase deviation .phi..sub.1 of the MR signals R from the moving part does not become zero, and the rephasing is not achieved.
As described above, in the prior-art MR imaging method, the polarity of the signal reading magnetic field Gr is reversed once to obtain the MR signals R. As a result, the phase deviation .phi..sub.1 of the moving part of the sample 1 does not become zero and rephasing is not achieved. Accordingly, artifacts occur and degrade the quality of the image.
Another example of prior-art system and a problem associated therewith will now be described with reference to FIG. 3. This prior art system is an example using the spin echo method. It is assumed that the RF pulse A1 is a 90.degree. pulse, the RF pulse A2 is a 180.degree. pulse, the NMR signal B is a spin echo signal, the X-axis gradient magnetic field Gr is a signal reading magnetic field for frequency encoding, the Y-axis gradient magnetic field Ge is a phase encoding magnetic field, and the Z-axis gradient magnetic field Gs is the slice magnetic field for designating the slice plane. It is also assumed that the NMR signals B are used for constructing an image by the two-dimensional Fourier transformation method.
First, a sample 1 is inserted in a static magnetic field generating unit 2, an RF coil 4, gradient magnetic field coils 9, 11 and 13, and the RF coil 4 and the Z-axis gradient magnetic field coils 13 are driven to simultaneously apply the RF pulse A1 and the slice magnetic field Gs1 to the sample 1. As a result, energy is supplied to the nuclear spins within the desired slice plane in the sample 1, and the phases of the nuclear spins begin to be disordered starting with the central position (peak time-point) of the RF pulse A1.
Next, the X-axis gradient magnetic field coil 9 and the Y-axis gradient magnetic field coil 11 are driven to apply the signal reading magnetic field Gr1 and the phase encoding magnetic field Ge, and to again apply the RF pulse A2 and the slice magnetic field Gs2. To align the spins in the direction of the slice (Z-axis) at the time of application of the slice magnetic field Gs1, the central position of the slice magnetic field Gs2 is shifted relative to the central position of the RF pulse A2.
After that, the NMR signal B is received while a signal reading magnetic field Gr2 having the same polarity as Gr1 is applied, and the NMR signal B is read in the computer 16. After the pulse area of the signal reading magnetic field Gr2 becomes equal to the pulse area of Gr1, i.e., upon expiration of the echo time TE from the central position of the RF pulse A1, the NMR signal B assumes its central position (peak value). Accordingly, the time-point at which the NMR signal B assumes its peak value is dependent on the time-points at which the RF pulse A1 and the signal reading magnetic field Gr2 are applied. The data of the NMR signal B are collected at a predetermined number of sample points while the signal reading magnetic field Gr2 is applied.
The above sequence is repeated a predetermined number of times corresponding to the predetermined number N (e.g., 256) of pixels while the phase encoding amount Ki determined from the pulse area of the phase encoding magnetic field Ge is varied at a predetermined pitch (see the broken lines). For example, when the number of pixels of the slice image is N.times.N, the number of the samples for the NMR signal B in a single cycle is not less than N, and the computer 16 performs two-dimensional Fourier transformation on the pulse series of the NMR signals B to reconstruct the slice image of the matrix size of N.times.N. This slice image is then displayed on the display unit 18.
In the above example, the signal reading magnetic field Gr1 is applied between the RF pulses A1 and A2. But it may alternatively be applied immediately before the signal reading magnetic field Gr2 with a reverse polarity. Moreover, a slice magnetic field Gs2 having its central position coinciding with the RF pulse A2 to align the spins in the direction of the slice may be applied immediately after the slice magnetic field Gs1 with a reverse polarity.
Now another prior-art method will be described with reference to FIG. 4 and FIG. 1.
FIG. 4 shows waveforms produced by means of the program of the computer 16 shown in FIG. 1 and output from the sequence controller 15. It is assumed that the two-dimensional Fourier transformation method is used in combination with the spin echo method.
First, a sample 1 is inserted in a static magnetic field generating device 2, an RF coil 4, gradient magnetic field coils 9, 11 and 13, and the RF coil 4 and the Z-axis gradient magnetic field coil 13 are driven simultaneously to apply an RF magnetic field pulse having a selected frequency (usually a 90.degree. pulse) RF1 and a slice magnetic field Gs1 for designating the slice plane. Energy is thereby supplied to the nuclear spins within the desired slice plane in the sample 1. The slice magnetic field Gs4 is applied so that the total area of Gs after the peak of RF1 is zero thereby to align the phase of the spins in the direction of the slice.
Then, the Y-axis gradient magnetic field coil 11 is driven to apply the phase encoding magnetic field Ge1.
Thereafter, a 180.degree. pulse is applied and then the X-axis gradient magnetic field coil 10 is driven so that, while the magnetic fields Gr8 and Gr10 are applied, an MR signal due to the spin echo is derived when the area of Gr10 becomes equal to the area of Gr8, i.e., TE after the peak of RF1. The MR signal is input to the computer 16 via the RF coil 4.
The time-point at which the MR signal assumes its peak is dependent on the time-point of the signal reading magnetic field and the RF magnetic field. The peak of 180.degree. pulse RF2 is made to appear at 1/2 between the peak of S1 and the peak of RF1.
The above-described MR imaging system is programmed to collect and image the signals of the stationary spins. Where there are moving spins, uncontrolled phase deviation and the introduction of spins from outside of the slice plan, artifacts may occur. Moreover, it is not possible to image the velocity, the flow velocity, the acceleration, the jerk, etc.
Another problem associated with the prior-art MR imaging method is that it does not take account of the movement of the sample, such as body movement of a patient, so that the resultant slice image contains artifacts.